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Vol. 83, Issue 6, 2123-2130, December 1997
,
1 Laboratorio di Ricerche
Cardiovascolar, ABSTRACT Castiglioni, P., R. Tommasini, M. Morpurgo, and M. Di
Rienzo. Modulation of pulmonary arterial input impedance during transition from inspiration to expiration. J. Appl.
Physiol. 83(6): 2123-2130, 1997.
pulmonary circulation; pulmonary blood pressure; pulmonary blood
flow; intrapleural pressure
INTRODUCTION THE PULMONARY ARTERIAL input impedance (Z), defined as
the ratio between the oscillations in blood pressure and blood flow at
the entrance of the pulmonary circulation, is a quantitative coupling
index between the right ventricle (RV) and the pulmonary vascular bed
in the frequency domain (12). Because of the low values of pulmonary
arterial pressure (PAP), the pulmonary input Z should be influenced by
the changes in intrathoracic pressure associated with breathing.
Surprisingly, the available literature does not report any significant
difference between the average Z values of heart beats occurring during
the inspiratory and expiratory phases of respiration, both at rest (3,
11) and during exercise (8). These studies, however, have not
considered the pulmonary Z during the transition between inspiration
(I) and expiration (E), namely, during the fraction of respiratory
cycle characterized by marked changes in intrathoracic pressure.
In the present study, we focused on this unexplored issue by evaluating
the pulmonary arterial input Z during the transition from I to E in
spontaneously breathing dogs. To interpret the results obtained by the
above analysis, we also separately investigated the influences of
respiration on blood pressure and blood flow, i.e., on the individual
determinants of the input Z.
METHODS
We investigated
whether respiration influences pulmonary arterial input impedance
during transition from inspiration to expiration in five anesthetized,
spontaneously breathing dogs. Impedance (Z) was separately assessed for
heart beats occurring in inspiration, in expiration, and during the
transition from inspiration to expiration (transitional beat).
Transitional beats were scored by the ratio between the fraction of
beat falling in expiration and the total beat duration
[expiratory fraction (Efr)] to quantify their
position within the transition. In transitional beats, input resistance
linearly increased with Efr; Z
modulus at the heart-rate frequency
(fHR) decreased up to
50% for Efr = 50%. Z phase at fHR was greater
than in inspiration for Efr <40% and lower for Efr >50%.
Unlike blood flow velocity, mean value and first harmonic of pulmonary
arterial pressure were correlated to
Efr and paralleled the changes of
input resistance and Z at fHR.
This indicates that respiration influences Z through modifications in
arterial pressure. The evidence of important respiratory influences on
Z function may help the pathophysiological interpretation of dysfunctions of the right heart pumping action, such as the so-called cor pulmonale.
1 · min1
infusion). Dogs were studied while they were in the right lateral position on a surgical table. They were intubated and then breathed spontaneously throughout the experiments.
Beat-by-beat estimation of Z. The preliminary step to estimate the input Z is the computation of the Fourier coefficients for each pair of PAP and FV waves. The Fourier coefficients of a single wave are mathematically defined as the coefficients of a periodic signal composed of infinite replicas of the original wave (see Ref. 10 for details). The possible occurrence of differences between start and end values of a single wave produces discontinuities in the periodic signal and may affect the accuracy of the estimation of Fourier coefficients because of the phenomenon known as aliasing. Before computing the Fourier estimates, we made a preliminary exploration of whether the characteristics of pressure and flow signals were compatible with the requirements of the Fourier analysis. The actual influence of aliasing on the spectral results was evaluated and was shown to be negligible on the first five harmonics of PAP and FV waves, namely on all the harmonics we considered in this study (see details in APPENDIX A). On the basis of this favorable validation, we computed the Fourier coefficients {Ak, Bk} for each PAP and FV wave. Equations B9 and B10 in APPENDIX B show the mathematical expression of Ak and Bk. From the Fourier coefficients, modulus Mk and phase
k of the
kth
harmonic were calculated for each wave by the formulas
Mk = (Ak2 + Bk2)1/2
and k = arctan(Bk/Ak),
with k = 1...5. The input Z modulus,
|Z(f)|, was computed at f = 0 Hz as the ratio between PAP and
FV mean values and at multiple frequencies of the instantaneous heart
rate (f = k × fHR) as the ratio between the
modulus of PAP and FV at the
kth harmonic.
Because Z was evaluated as the ratio between pressure and FV (instead
of flow volume), it was expressed in dynes per second per cubic
centimeters. The input Z phase, [Z(f)], was computed at
the same frequencies (f = k × fHR) as the difference between
the phases of PAP and FV
kth harmonics
(10).
Subsequently, four parameters were derived from each estimated
impedance curve Z(f): the modulus of Z(f) at f = 0 Hz [in short called input resistance
(RI)]; modulus and phase
of Z at the frequency of the instantaneous heart rate,
|Z(fHR)| and
[Z(fHR)],
respectively; and the characteristic Z modulus
(ZC), estimated by averaging the
Z modulus between 2 and 12 Hz, as suggested by Bergel and Milnor (1).
For each respiratory cycle, the differences between the estimates of
the above four parameters in the transitional beat and in the
inspiratory beat (the reference) were computed. The differences
RI,
|Z(fHR)|, and
ZC were then normalized with respect to the value in I to minimize variability between animals. Results of these four parameters were shown as a function of
Efr.
Heart rate, pressure, and flow during the transition.
To facilitate the interpretation of the results stemming from the
analysis of Z, we also evaluated the effects of the transition from I
to E on three factors that influence the Z estimation: heart rate, PAP,
and FV.
The reason why we investigated heart rate is apparent if one considers
that changes in Z(fHR) can be
due not only to a change in the Z curve, Z(f), but also to mere changes
in the frequency where Z(f) is evaluated, i.e., to changes in the
fHR. To estimate whether the
latter possibility occurred in our study, we computed the difference
between fHR in the transitional
beat and in reference to each respiratory cycle and plotted the
difference vs. Efr.
The separate behavior of PAP and FV during the transition from I to E
was also investigated to clarify the individual contribution of each
variable on the results obtained from the Z analysis. To this aim, we
considered the mean value and modulus and phase at the first harmonic
of PAP and FV. Modulations of these parameters were separately
quantified during transition by following the same procedure used for
the assessment of Z changes.
RESULTS
The results of the analysis on transitional beats for the whole group of animals are depicted in Figs. 3 and 4 as a function of Efr. RI progressively increased for augmenting Efr values, whereas |Z(fHR)| progressively decreased for Efr ranging from 0 to 50%, and thereafter symmetrically increased for Efr between 50 and 100%. At Efr values of ~50%, |Z(fHR)| was reduced to about one-half of the reference value on average. In no animal was this reduction less than
30%. The phase
[Z(fHR)] was positive with respect to the reference for
Efr <40% and became negative
for Efr >50%. In reference to
the ZC, we did not find any clear
dependence between ZC and
Efr (Fig. 4).
RI) and in impedance modulus
and phase at frequency of heart rate,
[|Z(fHR)|]
and [Z(fHR)], respectively, during transition from I to E. RI and
|Z(fHR)| are
expressed as %reference condition I. Data are plotted vs. Efr. 
, dog
1; 
, dog 2; 
,
dog 3; 
, dog
4; 
, dog 5.
ZC) during transition vs.
Efr. Data are expressed as
%reference condition I. Symbols are same as in Fig. 3.
Moreover, in the whole group of dogs, similar to the results obtained in the representative dog of Fig. 2, the Z phase of the transitional beats at harmonics 1, 2, 3, 4, and 5 was lower than in I for Efr between 10 and 90%, whereas the Z modulus was not influenced by Efr (data not shown). Because ZC was estimated as the average Z between 2 and 12 Hz, the latter finding also explains the lack of any dependence of ZC in relation to Efr. Heart rate, pressure, and flow during the transition. Variations of the fHR that occurred in transitional beats with respect to inspiratory beats are plotted vs. Efr in Fig. 5. In all animals, fHR was stable throughout the transition from I to E, and this result excludes the possibility that the observed changes in Z(fHR) might be due to changes in fHR.
fHR) during transition vs.
Efr. Data are expressed as
%reference condition I. Symbols are same as in Fig. 3.
Results of the separate analysis on PAP and FV waves are illustrated in Fig. 6. Figure 6 shows the relationship between the mean value and Efr and between the first harmonic modulus and phase and Efr. From a comparison of these data with the results of Fig. 3, it is apparent that the behaviors of mean value, modulus, and phase of PAP during transition parallel the behaviors of RI and of Z modulus and phase at fHR. In contrast, no evident modulation was observed for the FV-derived parameters as function of Efr, apart from a slight downward linear trend in the mean value. In the analysis of harmonics 1, 2, 3, 4, and 5, we found a decrease of PAP phase for Efr ranging between 10 and 90%, and no changes for PAP moduli, FV moduli, and FV phases (data not shown in Fig. 6).
PAP0), FV
(FV0), first harmonic modulus
(PAP1 and
FV1 ), and phase
([PAP1] and
[FV1]) during
transition vs. Efr for PAP and FV.
Data are expressed as %modulus in condition I or as difference from
phase in condition I. Symbols same as in Fig. 3.
DISCUSSION
A new procedure was developed for the beat-by-beat analysis of pulmonary arterial input Z as a function of the respiratory cycle. This allowed us to investigate specifically the changes of Z(f) during the transition from I to E.
Through application of our procedure, we observed that the transitional beats are characterized by a large variability in Z modulus and phase at the frequency of the heart rate, Z(fHR). This variability was proven not to be caused by changes of the fHR but rather by changes in the Z function [Z(f)]. These changes strictly depend on the onset time of the beat during the I-E transition. Specifically, the modulus of Z(fHR) displays a symmetrical pattern, which is characterized by a progressive drop from its reference value for Efr ranging from 0 to 50% and by an opposite progressive increase for Efr ranging from 50 to 100%. This has not been previously reported in literature. In particular, Murgo and Westerhof (11) measured PAP and FV in humans to compare input Z in I and E, but they classified transitional beats into one of the two phases of respiration according to where peak systolic FV occurred. They reached the conclusion that there was no difference between I and E in the overall Z spectrum. In a more recent study (3), our group investigated the influences of respiration on Z(f) by computing input Z in dogs during I, E, and postexpiration. Only beats completely occurring in a single respiratory phase were considered in that study. Significant differences were found between postexpiration and the other two respiratory phases but not between I and E, apart from a greater RI in E. In view of the symmetric nature of the relationship between |Z(fHR)| and Efr, it is now evident why in both these studies no change in Z modulus was detected between I and E.
Concerning the determinants of Z changes, our results indicate that the
changes in Z(f) are due almost entirely to changes in PAP, whereas
during the transition from I to E the modifications of mean FV are
negligible. It seems reasonable to ascribe the PAP changes and the
concomitant changes in Z(f) to the rapid and substantial compression of
the intrathoracic gases occurring between the end of I and the start of
E, which are in turn reflected by an increase in IPP. To obtain an
experimental support to this hypothesis, we also computed in two dogs
the input Z from the pulmonary blood pressure "purified" from the
influences of changes in IPP (Fig. 7). The
purified intravascular pressure was obtained by computing the
transmural pressure, namely PAP IPP, on the assumption that the
increase of IPP at the start of E induces an identical change in PAP.
This assumption is in line with previous observations indicating that
during the ventilatory cycle pulmonary intravascular pressure and IPP
undergo similar changes (4). Figure 8 shows
RI and
Z(fHR) obtained from PAP and
from PAP IPP in the two dogs. The removal of the influences of
IPP from PAP resulted in the abolition of the modulation of
Z(fHR) and in changes in the
linear relation existing between
Efr and
RI. These findings suggest that
IPP exerts a major role in the genesis of Z changes.
RI (left) and
Z(fHR) modulus
(middle) and phase (right) during transition vs.
Efr in 2 dogs before (open
symbols) and after (solid symbols) subtraction of IPP from PAP. Data
are expressed as %reference condition I. Symbols same as in Fig.
3.
On this basis, it remains to be explained how IPP may produce the
specific pattern observed in Z during the transition from I to E. A
possible explanation follows. The increase in IPP occurring during the
transition superimposes on the first harmonic of the purified PAP wave
a sinusoidal component with the same frequency but a different phase.
The difference between phases produces a lowering in the first harmonic
modulus of PAP that is more pronounced when the two components have
opposite phases, i.e., when the transition from I to E occurs in the
middle of the transitional beat
(Efr = 50%). A first support of
this reasoning is given by the experimental data of Fig.
9, which shows PAP, IPP, and transmural
pressure of three transitional beats, along with their respective first harmonic components, for Efr = 25, 51, and 80%. The maximal reduction of the PAP modulus occurs for
Efr = 51%, namely when the first harmonic components of IPP and transmural pressure are in counterphase.
Moreover, we theoretically verified whether the increase in IPP occurring during the transition from I to E was sufficient to explain the observed patterns in the mean value and first harmonic of PAP. This was achieved by developing the mathematical model described in APPENDIX B. The results of the simulation actually confirm that the IPP increase induced by the transition from I to E and the time of occurrence of this increase within the transitional beat (quantified by Efr) may account for all the changes observed in the mean value and in the first harmonic modulus and phase of the PAP wave.
As for the biological relevance of our findings, we should consider that in the pulmonary circulation the pulsatile component of pressure and flow is an important determinant of the RV afterload. Thus variations of Z(fHR) may substantially affect the RV work. The observed changes of Z(fHR) are therefore likely to influence the RV dynamics. Actually, alterations of RV systolic time intervals during the transition from I to E have been previously reported by our group (14), and the present findings suggest that Z(fHR) modulations may be one of the factors responsible for these alterations.
The potential clinical implications of the assessment of RV afterload through Z(f) have been recently illustrated in several papers. For instance, it has been suggested that changes in Z(f) may quantify the afterload increase that characterizes the so-called chronic cor pulmonale (5) and that Z(f) is a more sensitive indicator of vascular alterations than pulmonary vascular resistance for the assessment of pulmonary function of donor lungs before transplantation (2). Kussmaul et al. (7) showed that the ventricular-arterial coupling is importantly affected by pharmacological vasodilatation of pulmonary vessels in congestive heart failure and that the pulsatile properties of the pulmonary circulation must be taken into account to understand the effects of vasodilatation on cardiac output. More recently (4), pulmonary input Z has been used to evaluate the efficacy of nitric oxide administration and that of surgical interventions in the treatment and evaluation of chronic pulmonary hypertension. During the respiratory cycle, while shifting from I to E, the impairment in the RV afterload which characterizes pulmonary diseases may be further enhanced by alterations in the geometrical and mechanical characteristics of the proximal pulmonary arteries, as suggested for chronic cor pulmonale (9).
Moreover, the evaluation of the pulmonary input Z during respiration may be important for quantifying the milking-action effect, produced by the cyclical changes of lung volume due to respiration, on the blood circulation (5). This milking action is particularly important in the Fontan procedure when the RV is markedly hypoplastic.
In all these instances, the quantification of the ventricular-arterial coupling, as obtained by the estimation of the Z(f) during the respiratory cycle and, in particular, during the I-E transition, where we showed that important Z changes occur, may be of great clinical relevance.
ACKNOWLEDGEMENTS
We thank Prof. J. Milic-Emili (Montreal) for help in writing this paper.
FOOTNOTES
Deceased November 1995.
Address for reprint requests: P. Castiglioni, LaRC Centro di Bioingegneria, via Gozzadini 7, I-20148 Milan, Italy.
Received 23 December 1996; accepted in final form 4 August 1997.
APPENDIX A
Transitional Beats and Aliasing
The PAP and FV waveforms in the pulmonary circulation are influenced by the respiratory movements, and this may result in differences between the start and the end values of any single wave. When individual waveforms have to be analyzed by the Fourier series (as in the present study), a discrepancy between the onset and end of the waveform introduces a discontinuity that adds high-frequency components into the spectrum of the single waveform. Spectral components possibly added at frequencies higher than half the sampling rate are shifted toward lower frequencies and introduce distortion in the estimates. This error, known as "aliasing," can be avoided only by selecting a sufficiently high sampling rate (higher than two times the maximum frequency content of the signal). Because the discontinuity between starting and ending values results in an infinite number of harmonics, a residual aliasing error is unavoidable. Thus we quantified the practical influences of aliasing on our estimates of the Fourier components.The analysis was performed in two steps. First, we computed the
differences between starting and ending (head-tail) values of PAP and
FV waves of each transitional beat. This was done to quantify the
discontinuities actually occurred in our experimental data. Results are
shown in Fig. 10 where the
head-tail differences (dPAP and
dFV) are expressed as a
percentage of the wave amplitude for PAP and FV. The PAP differences
are close to 0% (almost no head-tail difference) for
Efr 
0, increase in a parabolic
fashion reaching a maximum (80%) for
Efr 50%, and progressively
return to 0% when Efr approaches
100%. No appreciable head-tail difference was found in the FV waves at
any value of Efr, thus excluding the possibility that aliasing may significantly affect the Fourier components of FV. On the basis of this finding, we continued the analysis by considering the effects of aliasing on PAP only.
In the second step of the analysis, we generated a simulated signal by summing 1) a real PAP wave characterized by the same initial and final values (this condition guarantees that the estimated Fourier components are not affected by aliasing) and 2) a sawtooth function f(t) that models possible head-tail discontinuities. This function was selected because its Fourier series is known in an analytic form (12); thus it is not affected by aliasing introduced by the estimation procedure. In particular, with W the amplitude and T the duration of the real PAP wave and dPAP defined as the head-tail difference of the simulated signal normalized by the wave amplitude W, then the saw-tooth function is given by the formula
|
(A1) |
|
(A2) |
It should be noted that the signal has an infinite number of harmonics produced by the discontinuities at t = nT (see Eq. A1). When the simulated signal is sampled at 200 Hz (i.e., at the same sampling frequency used in this study) components >100 Hz in f(t) are shifted toward a lower frequency band and cause aliasing.
Because the Fourier coefficients of f(t) are available in an analytic form and those of the real PAP wave can be estimated without aliasing, the true Fourier coefficients of the simulated signal are known. Therefore, we evaluated the effects of aliasing by comparing the theoretical Fourier coefficients of the simulated signal with those affected by aliasing that are estimated by the procedure used in our study.
The results, shown in Table 1, point out that, even considering discontinuities larger than those actually occurring in experimental data, the aliasing error in the estimation of the modulus of the first harmonic is only 5%. This indicates that reliable Fourier coefficients can be computed also in transitional beats characterized by major head-tail discontinuities in the PAP wave.
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APPENDIX B
A Mathematical Model of the IPP Influence on PAP at the Start of Expiration
In the text, we hypothesized that the changes in PAP during transitional beats might be due to the fast rise in IPP. In this APPENDIX, we show by means of a mathematical model that an increase of IPP resulting from the I to E transition is sufficient to explain the patterns observed in mean value and first harmonic of the PAP wave.Let us define
PAPI(t),
with 0 
t T, the PAP
waveform corresponding to a beat of duration T entirely
occurring during I, and
PAPIE(t)
the waveform corresponding to a transitional beat (assumed of the same
duration T). Moreover, let
IPP
(t) be the increase of IPP
during the transitional beat. Because of the fast pressure rise
occurring in IPP at the start of expiration, let us schematize IPP(t) with the following
stepwise function
|
(B1) |
|
(B2) |
(t), namely
|
(B3) |
, and
Eq. B3, we obtain the following
equations that link mean value and first harmonic modulus and phase of
the PAPIE waveform to the
corresponding Fourier coefficients of
PAPI and IPP
|
(B4) |
|
|
(B5) |
|
(B6) |
|
(B7) |
|
(B8) |
|
(B9) |
|
(B10) |
,
I
, and
as a function of the
Efr
|
(B11) |
![IPP<SUB>1</SUB>′(E<SUB>fr</SUB>) = <FR><NU>H</NU><DE>&pgr;</DE></FR> <RAD><RCD>2{1 − cos [2&pgr;(1 − E<SUB>fr</SUB>)]}</RCD></RAD>](/content/vol83/issue6/fulltext/2123/img018.gif)
|
(B12) |
![&phgr;<SUP>IPP</SUP><SUB>1</SUB>′(E<SUB>fr</SUB>) = arc tan <FR><NU>1 − cos [2&pgr;(1 − E<SUB>fr</SUB>)]</NU><DE>sin [2&pgr;(1 − E<SUB>fr</SUB>)]</DE></FR>](/content/vol83/issue6/fulltext/2123/img019.gif)
|
(B13) |
|
(B14) |
|
![= arc tan <FR><NU>PAP<SUP>I</SUP><SUB>1</SUB> sin (&phgr;<SUP>I</SUP><SUB>1</SUB>) + IPP<SUB>1</SUB>′ sin [&phgr;<SUP>IPP</SUP><SUB>1</SUB>′(E<SUB>fr</SUB>)]</NU><DE>PAP<SUP>I</SUP><SUB>1</SUB> cos (&phgr;<SUP>I</SUP><SUB>1</SUB>) + IPP<SUB>1</SUB>′ cos [&phgr;<SUP>IPP</SUP><SUB>1</SUB>′(E<SUB>fr</SUB>)]</DE></FR>](/content/vol83/issue6/fulltext/2123/img022.gif)
|
(B15) |
![PAP<SUP>IE</SUP><SUB>1</SUB>(E<SUB>fr</SUB>) = <FR><NU>PAP<SUP>I</SUP><SUB>1</SUB> cos (&phgr;<SUP>I</SUP><SUB>1</SUB>) + IPP<SUB>1</SUB>′ cos [&phgr;<SUP>IPP</SUP><SUB>1</SUB>′(E<SUB>fr</SUB>)]</NU><DE>cos [&phgr;<SUP>IE</SUP><SUB>1</SUB>(E<SUB>fr</SUB>)]</DE></FR>](/content/vol83/issue6/fulltext/2123/img023.gif)
|
(B16) |
PAPIE0,
, and
PAPIE1, as given by Eqs.
B14-B16, were computed for
Efr values ranging from 0 to 100%
to verify the ability of this model to describe the patterns observed
in our results. The values of the pressure parameters in I and the
increase of IPP during transition from I to E, as required to solve the
equations, were derived from the segment of experimental data
illustrated in Fig. 7 (dog 2). In
particular, PAPI0 was set equal to 29 mmHg, PAPI1 to 9 mmHg,
to 2.15 rad, and H to 9.8 mmHg. Figure 11 shows the
results of the simulation superimposed on the experimental results
obtained from dog 2. It is evident
that the model can faithfully reproduce
1) the linear trend in the mean
value of PAPIE for increasing
values of Efr;
2) the large reduction of the
PAPIE first harmonic for
Efr, close to 50%; and
3) the sigmoidal shape of the shift
between PAPIE and
PAPI phases that
characterize experimental data. Similarities between simulation and biological data corroborate the validity of the model
and provide a mathematical support to the hypothesis that most of the
changes observed in PAP are actually caused by the fast rise in IPP
occurring during the transition from I to E.
and 
. Right: changes of PAP phase at first
harmonic as estimated by model (dotted line) and observed in
dog 2 () vs.
Efr.
REFERENCES
| 1. |
Bergel, D. H.,
and
W. R. Milnor.
Pulmonary vascular impedance in the dog.
Circ. Res.
16:
401-415,
1965 |
| 2. |
Bittner, H. B.,
S. W. H. Kendall,
E. P. Chen,
D. Craig,
and
P. Van Trigt.
The effects of brain death on cardiopulmonary hemodynamics and pulmonary blood flow characteristics.
Chest
108:
1358-1363,
1995 |
| 3. | Castiglioni, P., M. Di Rienzo, M. Grossoni, R. Tommasini, and M. Morpurgo. Does spontaneous respiration alter pulmonary artery input impedance? Eur. Respir. J. 9: 2328-2334, 1996[Abstract]. |
| 4. | Chen, E. P., F. Tull, H. B. Bittner, D. Craig, R. D. Davis, and P. Van Trigt. Vascular impedance characterization in chronic pulmonary hypertension (Abstract). Chest 110, Suppl. 4: 29S, 1996. |
| 5. | Grant, B. J. B., and B. B. Lieber. Clinical significance of pulmonary arterial input impedance. Eur. Respir. J. 9: 2196-2199, 1996[Medline]. |
| 6. | Harris, P., and D. Heath. The Human Pulmonary Circulation (2nd ed.). Edinburgh, UK: Churchill Livingstone, 1977. |
| 7. |
Kussmaul, W. G.,
J. A. Altschuler,
W. H. Matthai,
and
W. K. Laskey.
Right ventricular-vascular interaction in congestive heart failure: importance of low-frequency impedance.
Circulation
88:
1010-1025,
1993 |
| 8. | Laskey, W. K., W. A. Ferrari, H. I. Palevsky, and W. G. Kussmaul. Pulmonary artery hemodynamics in primary pulmonary hypertension. J. Am. Coll. Card. 21: 406-412, 1993[Abstract]. |
| 9. | Malamani, V., R. Fogari, and M. Morpurgo. Effects of respiratory activity on the right ventricular pressure in chronic obstructive lung disease. Bull. Physiopathol. Respir. 8: 1437-1439, 1972. |
| 10. | Milnor, W. R. Hemodynamics (2nd ed.). Baltimore, MD: Williams & Wilkins, 1989. |
| 11. |
Murgo, J. F.,
and
N. Westerhof.
Input impedance of the pulmonary arterial system in normal man. Effects of respiration and comparison to systemic impedance.
Circ. Res.
54:
666-673,
1984 |
| 12. | Nichols, W. W., and M. F. O'Rourke. McDonald's Blood Flow in Arteries (3rd ed.). London: Arnold, 1990. |
| 13. | Press, W. H., S. A. Teukolsky, W. T. Vetterling, and B. P. Flann. Numerical recipes in Fortran. In: The Art of Scientific Computing (2nd ed.). Cambridge, UK: Cambridge Univ. Press, 1992. |
| 14. | Rustici, A., R. Fogari, M. Grossoni, R. Tommasini, G. Ukmar, and M. Morpurgo. Influence of spontaneous respiration on right ventricular systolic time intervals and pulmonary arterial pressures in the dog. In: Interaction Between Heart and Lung, edited by S. Daum. New York: Thieme Medical, 1989, p. 18-24. |
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ABSTRACT
